A. Field of the Invention
The present invention relates to methods and apparatus which employ X-ray or gamma ray radiography to obtain images of internal features of a human body, or objects within luggage or other such containers. More particularly, the invention relates to improved signal processing methods and apparatus for converting signals output from an array of X-ray radiation detectors which receive X-radiation emanating from an object irradiated by a source beam of X-ray radiation, to visually discernible images of internal features of the object.
B. Description of Background Art
Electromagnetic radiation of wavelengths substantially shorter than visible light, specifically X-rays and gamma rays having wavelengths less than about 0.1 nm, are routinely used to obtain visually discernible images of internal or sub-surface features of an object, by a method referred to as radiography. For example, transmission X-radiography has been in use since shortly after the discovery of X-rays to obtain visual images of internal features of human bodies, such as bones and organs. Transmission X-ray radiographic images are created by exposing an X-ray sensitive device to X-ray radiation which has been transmitted through an object from a source of X-ray radiation. Historically, photographic film plates were among the first X-ray sensitive devices used for X-ray imaging, and are still widely used in the medical field. In transmission X-ray radiography using photographic film plates, an X-ray source such as an X-ray tube emits a generally conically-shaped beam of X-ray radiation which irradiates a distant film plate holder. An object such as a human being or selected portion of the human's body is positioned between the X-ray source and the film plate holder, and upon being suitably positioned relative to the film plate holder and X-ray source, held stationary at that position. The X-ray source is then momentarily energized for a relative short, e.g., 1-second time period which has been calculated to be just sufficiently long to form an adequate image in the emulsion of a photographic plate held in the plate holder. The exposure time is kept as short as possible because X-radiation has a cumulative destructive affect on biological cells. Therefore, the dosage; which is equal to the product of X-ray intensity multiplied by exposure time, is desirably kept as small as possible.
Some X-ray-ray radiation which irradiates an object such as a human body is transmitted with little attenuation, while rays which impinge on denser internal portions of the body, such as bone, are more heavily absorbed or scattered, thus forming in the emulsion of the photographic plate a shadow image of the denser object features. The film plate is developed and fixed by conventional film processing chemistry reactions and is kept as a permanent visual record for viewing and analysis by medical professionals. One variation of transmission X-radiography, called X-ray fluoroscopy, utilizes in place of a film holder, a screen which visibly fluoresces in response to X-radiation, enabling real-time dynamic viewing of internal object features.
Another variation of transmission X-radiography utilizes in place of a film plate or fluorescent screen a matrix array of photodetectors which are overlain by a scintillator material that produces flashes of light or scintillations when irradiated by X-ray-radiation. Electrical signals output from the photodetectors are amplified and processed to form an electronic image of X-radiation incident upon the plane of the photodetectors. The electronic image can be converted to a visual image by a cathode ray tube (CRT) or Liquid Crystal Display (LCD) display device of a television or computer monitor. The electronic image can also be input to a computer which uses display recognition software to automatically recognize contraband such as guns and explosives hidden in luggage examined by an X-ray radiography system.
To reduce the dosage of radiation on an object, some newer X-radiography systems uses a collimator made of an X-ray absorbing material such as steel, which has a slit-shaped aperture which thus deforms a conically-shaped X-ray beam into a relatively thin, vertically elongated, wedge-shaped fan beam of X-ray radiation. Such systems utilize a vertically disposed linear array of one or more columns of scintillator-type X-ray detectors positioned at a target plane located on the far side of an object field positioned between the collimated X-ray source and the target plane. A mechanism is used to cause the fan beam of X-ray radiation to horizontally scan the entire width or horizontal extent of an object to be imaged. One method for causing the fan beam of X-ray radiation to scan an object utilizes horizontal motion of the object, on a conveyor belt for example, to move the object relative to a fixed X-ray radiation source and detector array. Another scan method used in Computerized Tomography (CT) scanning utilizes rotation of the X-ray source and, synchronous orbital motion of the collimator and detector plane, so that the beam remains at a fixed location on the detector array as the beam traverses the width of a stationary object located in the object field.
Whichever method is used to effect relative horizontal motion between a fan beam of X-ray radiation on an object, each instantaneous position of the beam on a particular column of detectors in an array produces detector output signals indicative of features in a single narrow vertically disposed slice of the object. Thus, as the fan beam of X-ray radiation horizontally traverses the object and detector array, a sequence of electrical signals is output from the detectors. This sequence of output signals corresponding to a vertical stack of locations of an object, must be concatenated to form a horizontal array of concatenated, sequentially sampled vertical signal slices to form a two-dimensional image of the object.
Now, the number of detectors in a linear array is generally relatively large, corresponding to the vertical extent of an object field to be scanned. Thus, for example, a typical linear detector array may consist of 1024 square detector elements, each having a length and width of 2 mm. Thus, because simultaneous parallel processing of 1024 detector signal channels would required unnecessarily large and complex electronic signal processing circuitry, most signal processing circuitry currently used on scanning X-ray radiography systems sequentially samples the outputs of a vertical detector array. This is accomplished by using a multiplexer to sequentially transfer or present output signals from individual detectors in the array to a circuit component such as an analog-to-digital converter (ADC). The time required to sequentially sample the output signals from each column or line of detectors corresponding to a single vertical image slice of an object, referred to herein as a scan line, must be shorter than the time required for the fan beam to traverse the object slice to an adjacent object-space slice, to ensure that no object feature is missed. Thus, for example, if the object is moved relative to the X-ray fan beam at a speed of 200 mm per second and the width of the detector array is 2 mm, scanning a line of detector image data must be completed in less than 2 mm/200 mm/sec= 1/100 second.
To accommodate requirements for different size object spaces, some scanning X-ray radiography systems in current use utilize a variable number of detector array boards, each having 2n detectors per board, where n can be any desired number, depending on the size of the detectors and the desired vertical extension or height of the detector plane, which is a function of the object field size and the distances between the detector plane, object space, and X-ray source. A plurality of detector boards may be arranged end-to-end to increase the number of detectors and height of the detector image plane.
Usually, each detector array board includes in addition to an array of detector elements some of the signal processing circuitry required for a scanning radiography system to form electronic images. Typical signal processing circuitry includes pre-amplifiers, and gain and offset adjustment circuitry. The present invention was conceived in response to certain limitations of existing scanning radiography systems, which will now be described.
A first type of prior art detector array board arrangement for use in scanning X-ray radiography systems circuitry in use utilizes one or more identical X-ray detector array boards, each board having 2n detectors. Typical numbers of detectors per board are 32, 64, 128, 256 or more. The first type of detector array board uses a single analog-to-digital converter (ADC). In this type board, a separate pre-amplifier is usually provided for each detector element in a detector array. Typically, each pre-amplifier includes a low-pass filter, and is sometimes referred to as a pseudo-integrator. The purpose of the low pass filter or pseudo-integrator is to develop an output voltage which is proportional to the time average of photodetector output currents produced in response to visible light scintillation photons which are produced in response to X-ray photons that irradiate the scintillator material, the visible photons in turn impinging on a photo-sensitive region of the photodetector. The low-pass transfer function of the pseudo-integrator is effective in averaging out statistical fluctuations in both the scintillator material photons and the photodetector output signal currents, thus reducing system noise.
In the first type detector array board for scanning digital radiography systems, the output terminal of each pre-amplifier or pseudo-integrator is connected to a separate input terminal of an analog multiplexer circuit. The output terminal of the multiplexer is in turn usually connected to an input terminal of an analog gain/offset compensation circuit, which in turn has an output terminal that is coupled through an analog switch to a board signal output terminal. Analog signals input by each detector pre-amplifier/pseudo-integrator circuit are sequentially transmitted through the multiplexer, under command of signals from a scan and transfer control circuitry which may be located on the board. Output signals from a single board or a plurality of detector array boards are sequentially input to a single ADC, which may be located exterior to the boards.
The first-type of detector array board for scanning X-ray radiography system requires a relatively fast ADC. Thus, a system with 1,000 detectors scanning at 200 lines per second requires an ADC conversion rate of 200 KHZ. Practical systems of this type in reality generally require a faster ADC, since time must be allowed for address decoding and settling time of signals on cables and buses. To reduce photon noise statistics without requiring the complexity and cost of utilizing true integrator circuits in each detector signal processing channel, each detector channel generally includes a pseudo-integrator (essentially a low-pass filter), which accumulates a new output signal that delays at a known rate to allow input and accumulation of a new signal of a new scan line. This circuit implementation requires only a single operational amplifier per detector channel. Selecting the time constant for the pseudo-integrators is a trade-off between signal-to-noise ratio and image sharpness. A shorter time constant will result in a sharper image but more photon statistic noise. Each detector array board of the first type is usually provided with gain and offset adjustment circuitry, which is used to tailor board operation to a particular system size, X-ray source type and strength, and system geometry.
Scanning X-ray radiography systems using a single ADC require transmission of analog signals through a cable/bus which may be many feet long; therefore, such systems are subject to noise interference from the environment. To reduce noise susceptibility, a second-type of prior art detector array board for scanning X-ray radiography systems currently in use utilizes a separate ADC for each board. Since placing a separate. ADC on each board increases the cost of each board, such boards often are provided with true integrators in place of pseudo-integrators. In this second, true integrator board, the signal output of each detector is captured by a separate analog integrator during each scan line. At the end of the scan line, charge from each integrator is transferred to a separate sample-and-hold circuit. Each integrator circuit is then reset by discharging its storage capacitor. During the next scan line, while new data is collected into the integrators, analog signals from each sample-and-hold circuit are input to an analog multiplexer and are sequentially output from the multiplexer through a gain/offset compensation circuit to an on-board ADC. Digital output signals from the ADC are transferred under the control of scan and transfer control circuitry to a host computer.
Since the second type, true-integrator board is provided with a separate, on-board ADC, analog signal paths are limited to a few centimeters, and confined to a shielded environment provided by an electrically conductive box which encloses components of the detector array board. The cost and complexity of each second type, true integrator detector array board is substantially greater than that of the first-type, pseudo-integrator board described above. This increase results from the larger number of components required for the true integrator circuits, sample-and-hold circuits, switching circuits, additional ADC and control logic. For example, each detector channel of the true-integrator detector array board requires 2 operational amplifiers and 2 analog switches. However, data output from the true integrator detector array board is less susceptible to photon statistic noise. Typically, the ADC is this type system samples each detector channel only once per scan line and passes converted data to an external host computer for image processing.
A third type of detector array board used in scanning X-ray radiography systems utilizes a plurality of n ADC's on each board, each ADC having an input terminal connected to the detector pre-amplifier output terminal of a separate detector channel. The digital output signal from each of the n ADC's is input to a register or memory storage location where it is accumulated to thereby effect digital integration. Digital integration boards of this type eliminate the requirement for all analog switching and signal processing functions, thus eliminating generation of spurious electronic noise. Such digital integrator boards generally use ADC's with relatively great resolution, e.g., 20 bits or higher, and produce a high quality image signal at high cost.
All prior art detector array boards for use in scanning X-ray radiography systems, including the three types described above produce output signals which must be further processed to produce electronic images of objects. Typically, the additional signal processing is performed in a data acquisition module located in a digital signal path between a host computer and detector array board or boards, but may optionally be performed solely within the host computer.
Each of the three detector array board types described above has certain less than optimum characteristics. For example, in a typical system which uses boards pseudo-integrators and a single ADC, the analog signal output from each pseudo-integrator of each detector channel is sampled and digitized once per scan line. In this implementation, selecting a time constant for the pseudo-integrator that is relatively long will result in low photon-statistic noise, but if the time constant is too long, electronic images derived from the signals will be blurred. A shorter pseudo-integrator time constant will result in a sharper image, but if made too short, variations in image brightness caused by photon statistic variations will become noticeable, particularly in low signal areas, and result in a grainy image. The foregoing constraints on pseudo-integrator time constants dictate that the time constant be reduced if the system scan rate must be increased, requiring that time-constant determining components of the pseudo-integrators be changed, which is usually only practicable by physically interchanging boards having components with different values.
In the second type, true-integrator detector array board described above, a practical limitation is imposed on board performance by a limitation in the amount of charge that each detector channel integrator circuit can hold. Using the same integrator circuits in a system with high power X-ray-ray sources or longer integration times can cause the charge storage capacitor of an integrator circuit to reach a maximum, saturated level while the detector output signal is still increasing, thus causing image information to be lost. To avoid integrator saturation resulting from higher detector output signals and/or slower scan times, it is necessary to change the value of the integrator storage capacitor to one having a higher value of capacitance. As a practical matter, this generally requires physically interchanging boards having the different components.
As mentioned above, the third type detector array board implementation which has a separate ADC for each detector channel avoids the limitations of the pseudo-integrator and true integrator boards described above, but is costly.
In view of the performance limitations and/or undesirably high cost of prior art detector array boards of the type described above, the present inventor developed a novel and improved system and methods for processing X-ray image data, which systems and methods are the subject matter of the present disclosure and described below.